Performance Evaluation of the Philips Gemini PET/CT System

Transcription

1 Performance Evaluation of the Philips Gemini PET/CT System Rebecca Gregory, Mike Partridge, Maggie A. Flower Joint Department of Physics, Institute of Cancer Research, Royal Marsden HS Foundation Trust, Downs Road, Sutton, Surrey, SM2 5PT, UK Abstract--Methods to test the practical performance of the Philips Gemini PET/CT system during clinical imaging are described and results presented. Measurements were performed using the EEC Emission phantom, alongside phantoms provided by Philips. The test methods used were based on EMA (U2 200) and those suggested by a task group of the Commission of European Communities (EEC). The clinical protocols in use were based on the single-pass emission-transmission whole body scanning protocols provided by Philips. Images are reconstructed using a 3D row action maximum likelihood algorithm (RAMLA). Attenuation correction was applied using both CT data (CTAC) and Cs-37 transmission scans (CsAC) to allow a comparison of the two methods. Images can be acquired with the PET and CT gantries adjacent (closed) or separated by 60 cm (open); geometric calibrations show that gantry alignment is most consistent in the closed position. Results show the non-uniformity of the images varied across the slices, for the non-attenuation corrected (OAC) images, but that uniformity was improved and more consistent in the both the CTAC and CsAC images. The reconstructed scatter fractions show that the RSF was better for OAC (0.5) than for CsAC (0.8) and CTAC (0.22), the sinogram derived scatter fraction was At the centre the transverse resolution was 5.3 mm in air (5.4 mm in water) and the axial resolution was 5.8 mm (6.5 mm). earer the FOV periphery the transverse radial resolution was 5.3 mm (5.3 mm), the transverse tangential resolution was 5.3 mm (5.5 mm) and the axial resolution was 7.5 mm (7.4 mm). The recovery coefficients show that partial volume only affects objects smaller than 22 mm. T I. ITRODUCTIO HE Philips Gemini PET/CT system consists of an Allegro 3D PET scanner in line with an Mx8000 EXP Dual-Slice CT Imaging System so that the patient can be imaged in both scanners on the same couch. The Gemini PET/CT system, at the Royal Marsden Hospital, has been in routine use for imaging cancer since in February A. Routine Acquisition Protocols The Allegro has no septa and operates exclusively in 3D mode. The Allegro scanner design has already been detailed in other publications []. Manuscript submitted October 5, A low dose CT scan (20kV, 80 mas/slice) is acquired during the patient study. The PET/CT scan allows for better localisation over single modality PET. The CT scan can be used for attenuation correction (CTAC). A Cs-37 source transmission scan is also acquired interleaved with the emission scan and stored using 4D rebinning (emission /transmission scan). This transmission scan is used for attenuation correction (Cs AC). B. Routine Reconstruction Protocols Currently the routine clinical images are reconstructed three times; without attenuation correction (AC), with CT AC and with Cs AC. All images are normalised to correct for the variation in detector efficiencies and distortion. All counts in each frame are decay corrected. The AC images use background subtraction which, from the level of background outside the body, estimates the background throughout the field of view and subtracts it, this background includes scatter and randoms that weren t removed using the energy window and delayed coincidence window respectively. To apply Cs AC the Cs-37 source transmission images are reconstructed using ordered subsets maximum estimated likelihood (OSEM). These images are corrected for emission contamination from the patient, scatter and then segmented. The segmented attenuation map is then forward projected into a sinogram of attenuation factors, which are then applied to the emission sinogram. CT AC is applied in the same way as in Cs AC, but the low dose CT image is used in place of the Cs-37 source transmission image and converted to PET attenuation coefficients. Because the CT image has better statistics, than the transmission image, it does not need to be segmented. All images are reconstructed routinely using 3D row-action maximum likelihood algorithm (RAMLA), a fast iterative algorithm [2], with a relaxation parameter of 0.006, for 2 iterations and a blob radius of 2.5. The image size is always 44 44, for a FOV of 576 mm, 4 mm thick slices. II. THE PERFORMACE TESTS Results of EMA tests, provided by the manufacturers, give an indication of the ideal performance of the system but not the performance in conditions for use during in vivo studies. The /04/$20.00 (C) 2004 IEEE

2 methods described within this paper aim to test the practical performance of the system during routine clinical imaging using locally available phantoms, and the routine clinical protocols. The test methods used are based on EMA U2 [3] and those suggested by a task group of the Commission of European Communities [4]. These tests have been performed using the EEC emission phantom. This phantom was designed to reflect the human shape, consisting of a whole-body phantom and a head-sized phantom, with various inserts that can be used to simulate small areas of activity uptake. The tests are designed to be reproducible, so that they can be repeated and initial results used as a base line for future performance checks. Uniformity and gantry alignment tests were performed using proprietary phantoms supplied by Philips. A. PET and CT Alignment Six 370 MBq a-22 point sources were carefully positioned using a jig provided by Philips. CT and PET images of the jig were acquired for the gantry in both the closed and open positions. The data were reconstructed, without AC and then analysed using software supplied by Philips, which measures the static offsets and rotation between the CT and PET image volumes. These offsets are stored, in the file header of subsequently acquired scans, and automatically applied to achieve image registration. Repeatability of these measurements was tested to see whether the changes may be due to the accuracy of the software or to the movement of the gantry between the open and closed positions. The alignment test was repeated 3 times with the gantry in the closed position. Then the test was repeated in the closed position and open positions, but moving the gantry between measurements. B. on-uniformity The uniformity phantom, supplied by Philips was filled with approximately 88 MBq of F-8. The phantom was then suspended from the end of the couch and then centred in the FOV. A low-dose CT scan was acquired followed by emission/ transmission scans. The images were reconstructed three times with each form of AC. on-uniformity is defined as the maximum relative deviation of counts from the mean. Positive and negative deviations were calculated, as + 00(max mu ) / mu and () + 00( mu min) / mu, where max is the maximum, min is the minimum and mu the mean of the counts, in a grid of small 2 mm square regions of interest (ROIs), defined within a circular useful region cm within the phantom, on each slice, as described in the EEC protocol [4]. C. Scatter Fractions The scatter source for the EEC phantom consists of a line source that can be positioned at 3 radial positions in the water filled cylindrical (head) phantom, at A=0, B=45 mm and C=90 mm. This scatter source was filled with approximately 4 MBq of Ga-68. At this activity the dead-time losses and the random coincidence rate were both below 5%. The head phantom was placed on the couch and centred in the field of view. Images were acquired with the scatter source at each of the radial positions. At each position a CT scan was acquired for attenuation correction (AC). Then emission/transmission scans were acquired for each source position. The PET protocol was modified to collect 0,000,000 counts per couch position. ) Reconstructed scatter fraction The images were reconstructed using the routine reconstruction algorithm, without AC, with CT AC and with Cs AC. The reconstructed scatter fraction was analysed using the EEC methodology [4]. In brief, the full width of half maximum (FWHM) of the source was determined by taking a profile across the image of the source. 3 ROIs were drawn on each slice of the image. ROI had a diameter of 4FWHM, and was centred on the image of the scatter source. The scatter within ROI was estimated from a concentric circular ROI (ROI 2 ), with diameter 4 FWHW plus 2 pixel widths. An ROI with the diameter of the head phantom was drawn centred on the source at the central position (ROI 3 ). These ROIs are shown in the diagram, Fig.. The reconstructed scatter fraction, RSFi, the scattered counts S i divided total counts, scattered plus true, T i, was calculated for each slice i and source position. RSFi = Si /( Si + Ti ) = ( C3 ( C A )) / C (2) 3 where C 3 is the counts in ROI 3, there for the total scattered plus true counts. C is the number of counts in ROI. A is the estimate of the scattered counts in ROI, calculated as, A ( ) ( /( ) = C2 C 2 ) (3) where C 2 is the counts in ROI 2, and 2 is the number of pixels in ROI 2 and is the number of pixels in ROI. These scatter fractions describe the scatter behaviour at each radial position of the source. ROI Source ROI 3 ROI 2 Fig.. A schematic diagram showing the position of the regions of interest for the source positioned at the centre of the cylinder. To give the average scatter fraction for the whole of the uniform cylinder, RSF cyl, the individual RSF for each source position are weighted and summed. The weighting corresponds to the fraction of the area of the cross-section that they present /04/$20.00 (C) 2004 IEEE

3 S Ai + 8 S Bi SC i = (4) RSFcyl C3Ai + 8 C3Bi C3C i where S Ai, S Bi, S Ci are the scattered counts in each slice for the source at positions A, B and C respectively, these are summed over slice. C 3Ai, C 3Bi and C 3Ci are the counts in ROI 3 in each slice for the source at positions A, B and C respectively. 2) Sinogram derived Scatter fraction For comparison, the scatter fraction was also calculated from the interpolated sinogram, for the line source at position B, which was rebinned into 2 mm slices. This is based on the test described in [3]. For each projection angle, the peak was shifted to align with the central pixel. Then for each slice, the counts within ±20 mm of this peak were summed over all the projections. For each peak profile, the counts below a line joining the counts at the edge of the 40 mm strip (C L and C R ) are considered to be scattered and the counts in the peak above to be unscattered, see Fig. 2. Therefore the scatter fraction for each slice i, SF i, is given by, 0.5p( CR + CL ) + C0 SFi = (5) CT where C T is the total counts under the peak, p is the number of pixels between the edges of the strip and C 0 the number of counts outside the strip. The system scatter fraction is weighted average of the slice scatter fractions. C L Unscattered counts Scattered counts C R -20mm 0mm +20mm Fig. 2. Illustration of scattered and un-scattered counts. D. Transaxial and Axial Resolution Two line-sources, consisting of glass tubes filled with F-8 were placed in the head phantom. One was aligned along the central axis and the other parallel and 9 cm away. The central axis of the head phantom was aligned with the z-axis and rotated so that images could be acquired with the second line source displaced along the x-axis and then displaced along the y-axis. Then the central axis of the head phantom was positioned along the x-axis and the 2 line-sources; one positioned at the centre and the other 4.5 cm away in the z- direction, so that the axial resolution could be measured. These measurements were then repeated for the head phantom both empty and filled with water. PET images were acquired over a small FOV of 256 mm and the reconstruction protocol was amended for a matrix size of to give a pixel size of mm (less than /3 full width half maximum (FWHM)). The resolution was calculated as the average tangential and axial FWHM s of a profile taken through and perpendicular to the images of the line sources, following the U 2 guidelines [3]. E. Recovery Coefficients EEC spheres, placed in the water-filled head phantom, were used to measure the recovery coefficients. All the spheres were filled with Ga-68, from the same stock vial, with an activity concentration of 0.57 MBq/ml. Another head phantom was filled with Ga-68 to give an activity concentration of MBq/ml. A CT scan of the head phantom with the spheres was acquired. To ensure that the spheres are centred on a reconstructed image slice, three emission and Cs-source transmission scans were acquired of the spheres, at axial intervals of 0.5 times the reconstructed slice width. A CT scan was then acquired of the head phantom, without the spheres, followed by emission and Cs-source transmission scans. The scans were reconstructed, using the routinely used clinical reconstruction protocol, with CT AC and Cs AC. The images of the spheres were viewed to find the axial position where the smallest sphere was brightest; the scan at this position was then used for the following analysis. Recovery coefficients were measured for both types of AC. Small regions of interest (ROIs), roughly the same diameter as the resolution of the system (2 mm diameter) were placed on the image of each of the spheres. And a large ROI (50 mm diameter) was placed on the image of the head phantom, without the spheres, on a slice at the same position as the spheres. The recovery coefficient, R Si, for each sphere was calculated using, R Si = ( CSi As ) ( CH AH ), (6) where C Si and C H are the count rate per second per pixel, in sphere i and the head phantom respectively. A S and A H are the activity concentrations in the spheres and the head phantom respectively, decay corrected for the time of each scan. III. RESULTS A. PET and CT Alignment Initial results, showing the variation in static offsets with time for both the open and closed gantry positions, are shown in Fig. 3. The results of the reproducibility tests are shown in table I. TABLE I RESULTS OF THE ALIGMET REPEATABILITY MEASUREMETS /04/$20.00 (C) 2004 IEEE

4 The coefficient of variation of the non-uniformity of the ROIs, defined on each slice are also plotted to demonstrate how the uniformity varies within, the useful region on each of the slices. Due to set up errors the last 2 cm of the CT attenuation corrected images are missing. C. Scatter Fractions ) Reconstructed scatter fraction The variation in reconstructed scatter fraction across the slices is shown in Fig. 5, for the on AC image. There is an apparent increase in the scatter fraction toward the 90 mm slice position, this is due to the presence of the Cs-source (which is stored at that end of the scanner), and is not scatter. In order to avoid this area the volume scatter fractions were calculated using the central values in the flatter region of the graph (from 65 to +65 mm). The volume reconstructed scatter fractions were 0.5, 0.8 and 0.22 for the non-ac, Cs AC and CT AC images respectively. Figure 3. Graphs to show the variation in static offsets since the PET/CT scanner went into clinical use. Graph a shows the values for the gantry in the closed position, b) shows the values for the gantry in the open position and c) the rotation about the z-axis for both the closed and open gantry position (rotation about the x and y axis is not determined) B. on-uniformity Results of the non-uniformity measurements are shown in Fig. 4. These are displayed as the maximum positive and negative deviations in each of the slices, to give an indication of how the maximum non-uniformity changes across the slices. Fig. 5. Reconstructed scatter fractions versus slice for each source position for the non-ac image. Fig. 4. Results of the non-uniformity test, on the top row the range of nonuniformity values measured in 45 slices, and on the bottom row the coefficient of variation of non-uniformity values for each slice, a) for non-ac, b) for CT AC and c) for Cs AC /04/$20.00 (C) 2004 IEEE

5 2) Sinogram derived Scatter fraction The volume scatter fraction, calculated in the area unaffected by the transmission source was 0.36, the variation in scatter fraction across the slices is shown in Fig. 6. Fig.6 A graph of scatter fractions (calculated from the sinogram) vs. slice position. D. Transaxial and Axial Resolution The results of the resolution measurements are shown in table II. TABLE. U-2 RESOLUTIO E. Recovery Coefficients The recovery coefficients were plotted against the sphere diameters, these are shown in Fig. 7. As expected these values tend towards.0, a double exponential, tending to was fitted to the data (Trend line). Recovery Coefficient y = ( 5.5e 'CS AC' 'CT AC' Trend line Sphere Diameter (mm) Fig. 7. Recovery coefficients as a function of sphere diameter. IV. DISCUSSIO )( 0.9e 0.2x 0.7 x A major advantage of a PET/CT system is the simplification of image registration. Alignment of the images assumes that the position of the PET system relative to the CT system is known. Initial checks show that the PET and CT scanner alignment needs to be calibrated regularly, but over time the variation in alignment has settled, as the PET/CT system and building have settled. The reproducibility measurements show ) that the variation in offsets is not due to uncertainty in the measurements, and that moving the gantry does not affect the alignment of the inherently registered images. on-uniformity measurements tested the PET scanner's response to a homogeneous activity distribution. The EEC protocol [4] states that non-uniformity measurements should be made on the EEC head phantom. However, Philips provide a cylinder to be used for monthly uniformity checks, so this cylindrical phantom was used for consistency. These results show that, there is more non-uniformity towards the end slices for the non-ac images, than the AC images. The degree of non-uniformity is comparable to the GE Advance nonuniformity [5]. Scattering of photons, within the phantom and off the gantry components, causes coincident events to be incorrectly localised. The reconstructed scatter fractions show that the attenuation corrected images are more sensitive to scatter effects than images without attenuation correction. The CT AC images are effected more that the Cs AC images. The intrinsic scatter fraction derived from the sinogram is very different, but this is due to the different analysis technique. The spatial resolution measured in both air and water are very similar, the scatter of radiation in the water has caused very slightly larger FWHMs in some cases, but the difference is very small. As expected the transverse resolution is better than the axial resolution. Over the region measured, there was no apparent degradation in spatial resolution for sources away from the central axis, due to the energy absorbed in the scintillator being spread over a greater area, and therefore over more detector elements. The results of the recovery coefficient measurements quantify the apparent decrease in tracer concentration in an ROI, when the dimensions of the object in any direction are less than 3 FWHM. As the average FWHM is about 5.5 mm in water, it may be expected that the apparent tracer concentration will not recover until the object is over 6.5 mm in all dimensions, therefore full recovery would only occur after the 3rd data points, this is observed in the graph of recovery coefficients. The recovery coefficients are fairly similar for both forms of attenuation correction. V. REFERECES [] S. Surti and J.S. Karp, Imaging characteristics of a 3-Dimensional GSO whole-body PET camera J ucl Med., vol. 45, no. 6, pp , Jan [2] J. Browne, A row-action alternative to the EM algorithm for maximum likelihoods in emission tomography IEEE Trans. Med. Imag., vol. 5, no. 5, pp , Oct [3] EMA U 2-200, Performance measurements of positron emission tomographs, EMA standards publication 200. [4] R. Guzzardi, C.R. Bellina, B. Knoop, K. Jordan, H. Ostertag, H.W. Reist et al., Methodologies for perfomance evaluation of positron emission tomographs J. ucl. Biol. Med. Vol. 35, pp. 4-57, July 99. [5] T.K. Lewellen, S.G. Kohlmyer, R.S. Miyaoka, M.S. Kaplan, C.W. Staerns, S.F. Schubert, Investigation of the performance of the General Electric ADVACE positron emission eomograph in 3D mode. IEEE Trans. ucl. Sci, Vol. 43, o. 4, pp , 996, Aug /04/$20.00 (C) 2004 IEEE